Republic of Iraq Ministry of Higher Education and Scientific Research Al-Nahrain University College of Science Department of Physics Surface Modification of Titanium and Titanium Alloy Using Ceramic Biomaterials by RF Sputtering A Dissertation Submitted to the College of Science / Al-Nahrain University as a Partial Fulfillment of the Requirements for the Degree of Doctorate of Philosophy in Physics By Dunya Abdulsahib Hashim Hamdi B.Sc. Physics/College of Science / Al-Nahrain University (1997) M.Sc. Physics/College of Science / Al-Nahrain University (2003) Supervised by Dr. Thamir A. Jabbar Jumah Dr. Thair Latif Al-Zubaydi (Assistant Professor ) (Senior Scientific Researcher) March 2016 A.D Jamadi Al- Akhira 1437A.H
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Alloy Using Ceramic Biomaterials
Submitted to the College of Science / Al-Nahrain University as
a
Partial Fulfillment of the Requirements for the Degree of
Doctorate
of Philosophy in Physics
Supervised by
(Assistant Professor ) (Senior Scientific Researcher)
March 2016 A.D Jamadi Al- Akhira 1437A.H
DEDICATION
Mother & Father
this study…
Dunya
Acknowledgement
First of all I want to thank God Almighty for providing me
with
willingness and strength to accomplish this work and I pray that
his blessings
upon me may continuous throughout my life.
Great thanks and deepest gratitude to my supervisors Dr.
Thamir
A.Jabbar and Dr.Thair Latif Al-Zubaydi for their valuable
scientific guidance,
inspiring discussions, continuous support and great efforts in
supervision this
study.
I am gratefully thankful to Dr. Ayad Murad, Administrative
Assistant to
the President of the University, for his continuous help, guidance
and support
throughout the all research period.
Many thanks go to Dr. Hadi Mohemmed Ali, Dean of College of
Science,
University of Al Nahrain for his support to the higher studies
program.
I am gratefully thankful to Dr. Alaa Jabbar Ghazai, Head of
physics
Department, for his constant care, valuable advice and continuous
support
throughout the course of the study.
My sincere thanks to my supervisors Dr. Zhong Tao , Murdoch
University
Australia and Dr. No , KAIST University Korea for her interest work
of this
study and their never-ending support and help whenever needed
throughout my
mission scholastic.
Deep appreciation to Mr Akeel F. Hassan and Mr. Jammal F.
Hamodi
for their great help that I will never forget
Dunya
I
Summary
The current in this thesis research deals with the plasma
sputtering which was
used to deposit single and triple layers ceramics on (cpTi) and (
Ti-6Al-4V)
alloy. Also ,in this work the in vitro electrochemical tests ,
formation of HAp
by biomimetic process on the coated layers by the immersion in
Simulated
Body Fluid(SBF) at PH equal to 7.4 and room temperature was done .
The
single layer TiO2and HAp , multi-layer consist of TiO2 and Al2O3
were
deposited on cpTi and Ti-6Al-4Valloy, followed by the deposition of
third
layer of HAp. The deposition conditions in the RF sputtering system
such as
vacuum pressure, substrate temperature, power, gas type, gas flow
and
deposition time were fixed for the sputtering system where as the
distance
between the target and the substrate was varied (9 and 7 cm).
Structural analysis
was carried out , on the deposited layers(single and triple) using
X-ray
Diffraction (XRD), Scanning Electron Microscopy (SEM),
Furrier
Transformation Infra-Red (FTIR), Raman Spectroscopy and X-ray
Photoelectron Spectroscopy (XPS). Elemental analysis for HAp
deposited in the
single and triple layer was done using Energy Dispersive X-ray
Spectroscopy
(EDS).
To investigate the biocompatibility of coatings the formed layers,
were
immersed in (SBF) for one month. After one month the samples were
taken and
anlyezed using the same structural analysis techniques used before
immersion.
Electrochemical investigation was carried out on the deposited
layers used
Open Circuit Potential (OCP),Linear Sweep Voltage(LSV) and
Electrochemical
Impedance Spectroscopy(EIS) in SBF. The XRD structural results show
that
the formation of HAp which was dominated on the phase formation on
the
surface of the single and triple layers. Also another Calcium
–Phosphorus
compound phases are found such as Tri Calcium Phosphate(TCP),
Octa
Calcium Phosphate(OCP) and Calcium Phosphate(CP) along with small
fraction
II
of the Titanium phase belong the substrate .For TiO2 single layer,
XRD pat reins
shows the formation of the rutile TiO2which covered the substrate
surface.
Elemental analysis by using EDS for the single and triple layer
deposited in
cpTi and Ti-6Al-4Valloy shows the energy transitions belong to the
dominated
elements, Ca and P, which contained in the HAp layer that already
showed by
XRD. Surface analysis by XPS for immersed samples in SBF for one
month
show the bonding type and the compound that formed. The XPS
results
confirmed the SEM observation, the FTIR and Raman results. It was
found
from the XPS analysis that most of the compound covers the surface
are
belongs to the Ca-P companies to the carboxyl groups (C-O, C-H) and
this was
confirmed by the XRD investigation which show the domination of
(HAp) in
the highest intensity (211) reflection at 2θ of 31.7 for the single
and triple layer
coating. The XRD results also show the disappearance of the other
CaP
compounds after one month of immersion in SBF like TCP and CP
.The
increase in the HAp concentration after immersion indicates the
occurrence of
the biomimetic process that increases the precipitation of Ca + and
P
- from the
SBF and the two roots combined together to form CaP compound that
increases
the HAp concentration. The electrochemical parameters predicted
from
corrosion test show the improvement in the corrosion resistance of
both cpTi
and Ti-6Al-4Valloy after coated by shifting the OCP toward the
nobel direction
which was the same behaviour for the corrosion potential Ecorr, and
decreases in
the corrosion current Icorr and the corrosion rate CR comparing to
the uncoated
specimens. The (EIS) measurements conformed the improvement in
the
corrosion parameters result of from OCP and LSV by the very low
capacitance
for the coated specimens compared to that for the uncoated one
which means
that the single layer and triple layer protect the surface by
increasing the
equivalent circuit impedance blocking the path way for the active
ions like Cl -
and O + to attack the base metal.
III
Replacement Materials
2.3.1 Commercial Pure Titanium cpTi 21
2.3.2 Titanium Aluminum Vanadium Ti-6Al-4V
alloy
21
2.6 Ceramic 30
2.8 Corrosion 38
2.8.2 Electrochemical Kinetics of Crrosion 40
2.9 Corrosion Test Methods 41
2.9.1 Open Circuit potential (OCP) 41
2.9.2 Linear Swap Voltage (LSV) 42
2.9.3 Electrochemical impedance spectroscopy
IV
2.13 Corrosion of Biomedical Metals and Alloys 53
CHAPTER THREE
Target (Pilot study)
Instrument
60
3.2.6 In vitro Electrochemical Reaction Test 61
3.2.7 Experimental Works includesTiO2,Al2O3 and
HAp as Targets (Research study)
62
3.3.3.1 Single Layer HAp&TiO2 69
3.3.3.2 Multilayer(TiO2+Al2O3) 69
3.3.4 Heat Treatments of Samples 70
3.3.5 Surface Examination and Characterization
Instrument
72
(EDS)
72
3.3.5.3 Phase Analysis 74
3.3.5.5 Fourier Transform Infrared Spectroscopy
(FTIR)
77
3.3.8 In vitro Electrochemical Tests 83
V
study)
82
4.1.2 XRD Single Layer High Thickness Film TiO2 83
4.1.3 Morphology of Thin Film TiO2 84
4.1.4 In vitro Tests Coating Morphology of Thin
Film TiO2
and HAp(Research study)
cpTi and Ti-6Al-4V Alloy
4.2.1.4 EDS of Triple Layer HAp 93
4.2.2 Structural Analysis 96
4.2.3 Thickness and Morphology of Thin Film
Using SEM
4.2.3.2 Morphology of Layer Thin Film HAp 107
4.2.4 Elemental Depth Profile at TiO2 Layer Using
XPS
112
Triple Layers
4.2.8 Biomimetic Immersion Test 124
4.2.8.1 EDS Analysis of HAp Triple Layer After
Immersion
124
Immersion
126
Immersion
127
VI
Immersion
133
chemical study)
4.2.9.2 LSV-Linear Sweep Voltammetry 136
4.2.9.3 EIS Electrochemical Impedance Statistical
Measurements
139
Appendix B Ti 6Al 4V Alloy Certificate B1
Appendix C TOBAS Program Final Report of XRD
Lab. for Apatite Concentration
2.2 Some of the physical properties of anatase and rutile
TiO2 structures at room temperature
36
3.1 The TiO2 target coated on substrates (cpTi and Ti-6Al-
4V) alloys as single layer
58
standard PDF files
3.4 The condition of cpTi and Ti-6Al-4Valloy coated with
single layer HAp
triple layer
with their Surface modification types
81
4.2 Corrosion parameters for cpTi and Ti-6Al-4V alloy
uncoated and coated with different thickness TiO2
88
determined by EDS
4.4 Concentrations single layer HAp for tD (2, 4, 7
and10) hours coated on cpTi and Ti-6Al-4V alloy
determined by EDS
91
4.5 Concentrations single layer HAp for tD (2, 4, 7 and10)
hours , D9 and D7 respactively coated on cpTi and
Ti-6Al-4V alloy determined by EDS
93
4.6 Concentrations triple layer HAp for tD 10 hours 95
VIII
EDS
4.744 4.7 The particle size HAp film for tD 7 hours coating
cpTi
& Ti-6Al-4V alloy .
4.8 The concentration of apatite in single and triple layer
coating cpTi- Ti-6Al-4V alloy
4.9 Thickness of film coated on cpTi and Ti-6Al-4V alloy
as single and triple layers.
107
4.11 XPS spectrum analysis high resolution of C1s, O1s
,Ca2p and P2p of HA film coating cpTi and Ti-6Al-
4V alloy as single and triple layers.
123
4.12 EDS data results of immersion Ti-6Al-4V alloy coated
with HAp for tD 10 hours as single and tribe layer in
SBF
125
,Ca2p and P2p coated Ti-6Al-4V alloy with triple
layer of HAp immersed in SBF for one month
133
uncoated and coated with HAp as single and triple
layers
138
and triple layers
20
2.4 Schematic representation of diode sputtering assembly 28
2.5 Schematic illustration of the development of a negative
bias in a RF system : (a) The current-voltage
characteristic for an electrode immersed in plasma,(b)
When an alternating voltage is applied to an electrode, a
positive/negative potential appears on the surface
29
together forming
30
hexagonal
34
36
2.10 Experimentally measurable anodic and cathodic
polarization curves
circuit
45
versus logarithm of current density for concentration
polarization
activation-concentration polarization
for AC and R
2.15 Crevice Corrosion 52
2.16 Pitting Corrosion 52
c: piston hydraulic, d: target before sputtering
e: target after sputtering
3.3 plane diagram for the experimental works (research
study) and (b) diagram for RF sputtering system
64
3.4 Grinding paper and mission 65
3.5 (a) and (b) the ultrasonic etching and clean for the
substrates.
65
3.7 (a) Sputtering system, (b) Substrate chamber,
(c) and (d) Sputtering chamber and (e) Targets positron
68
3.9 Energy Dispersive X-ray Spectroscopy (EDS) 72
3.10 (a) Scanning Electron Microscopy (SEM) ,(b) sputter
coater device and (c ) holder samples
73
3.12 ( a) XPS system ;(b) Sample inside system and(c)
System screen to select the analyzing point on the
sample
76
3.14 RAMAN Scattering Instrument 78
3.15 (a) and (b) Alpha Optical Step 79
3.16 The samplesTi-6Al-4V were immersed in SBF for one
month after coated with (a)HAp as single layer
,(b)(TiO2+Al2O3+HAp)
81
XI
coating with TiO2 film
4.2 XRD patterns of cpTi alloy coated with TiO2 with
thickness 500nm.
84
4.3 Top viewSEM (a) cpTi uncoated,( b )and( c) film TiO2
with thickness 500nm coated on surface of cpTi and Ti-
6Al-4V alloy respectively
4.4 LSV pattrens for the uncoated and TiO2 coated cpTi
with different thickness in SBF solution
87
4.5 LSV pattrens for the uncoated and TiO2 coated Ti-6Al-
4V alloy with different thickness in SBF solution
87
and (d) triple layers (TiO2+Al2O3+HAp)
88
4.7 EDS spectra for (a) cpTi and (b) Ti-6Al-4V alloy un
coated
89
4.8 EDSspectra of single layer HAp for tD 10 hours an D9
coated on (a) cpTi and (b)Ti-6Al-4V alloy
91
4.9 EDS spectra of triple layer HAp for tD 7 hours and
D7coated on (a) cpTi and (b) Ti-6Al-4V alloy .
92
4.10 EDS spectra of triple layer HAp for tD 7 hours and
D7coated on (a) cpTi and (b) Ti-6Al-4V alloy
95
4.11 XRD patterns of the uncoated and HAp coated cpTi
alloy for different deposition time tD(single layer) low
thickness
97
4.12 XRD patterns of the uncoated and HAp coated Ti -6Al-
4Valloy for different time deposition time tD(single
layer) low thickness
4.13 XRD patterns of the uncoated and HAp coated cpTi
alloy for different deposition time tD(single layer) high
thickness
99
4.14 XRD patterns of the uncoated and HAp coated Ti -6Al-
4Valloy for different deposition time tD(single layer)
high thickness
99
XII
4.15 XRD patterns of the uncoated and HAp coated cpTi for
different deposition time tD(triple layer) high thickness
101
4.16 XRD patterns of the uncoated and HAp coated Ti -6Al-
4Valloy for different deposition time tD(triple layer) high
thickness
101
4.17 Cross section SEM micrograph of HA coated on Ti-
6Al-4V alloy :
(a) at D9 cm distance as single layer for tD 10 hours,(b)
at D7 cm distance as single layer for tD 10 hours ,(c) at
D7 cm distance as triple layer tD 4 hours,(d) at D7 cm
distance as triple layer tD 7 hours
106
4.18 Top-view SEM image of HA as single layer for tD10
hours coated on:(a) cpTi at D9cm ,( b) Ti-6Al-
4Valloy at D9 cm,(c) cpTi at D7 cm , (d) Ti-6Al-
4Valloy at D7 cm
Al2O3and coated on:
( a) cpTi ,(b) Ti-6Al-4Valloy at D7 cm and tD 4 hours
110
4.20 Top-view SEM image of HAp coated (a) cpTi and (b)
Ti-6Al-4Valloy as triple layer at D7 cm for tD 7 hours
111
4.21 Atomic percent ves. Etch Time of TiO2 layer coated Ti-
6Al-4V alloy
112
4.22 Typical FTIR spectrum of HAp coated (a) cpTi and (b)
Ti-6Al-4Valloy as triple layer at D 7 cm for tD 10 hours
115
4.23 Typical RAMAN spectrum of HAp coated (a) cpTi and
(b) Ti-6Al-4Valloy as triple layer at D7 cm for tD 10
hours
116
4.24 XPS spectrum survey scan of thin film HA with
precipitated apatite coated on Ti-6Al-4V alloy for single
and triple layer at D7 cm for tD 10and 7 hours
respactively
117
4.25 XPS spectrum high resolution of C1s, O1s ,Ca2p and
P2p thin film HAp with precipitated apatite coated on
cpTi and Ti-6Al-4V alloy for single and triple layer at
D7 cm for tD 10 hours
122
after coated with :(a) HAp as single layer and (b)three
layer (TiO2+Al2O3+HAp)
4.27 EDS results of immersion Ti-6Al-4V alloy coated with
HAp tD 10 hours as (a) single and (b) triple layer in SBF
125
4.28 XRD patterns of Ti-6Al-4V alloy coated with HAp for
tD 10 hours as (a) single layer and (b) tribe layer after
immersion in SBF for one month
126
immersion in SBF for one month
128
4Valloy coated by film HAp as triple layer after
immersion in SBF for one month
129
4.31 XPS spectrum survey scan of immersion Ti-6Al-4V
alloy coated by thin film HAp as triple layer at D7 cm
and tD 10 hours .
of HAp immersed in SBF for one month .
132
4.33 The steady-state OCP for cpTi coated with HAp single
and three layers in comparsion to uncoated in simulated
body fluid .
in simulated body fluid
4.35 LSV diagrams for cpTi coated with HAp single and
three layers in comparsion to uncoated in simulated
body fluid
simulated body fluid
4.37 Nyquist plots of the uncoated and single layer HAp
coated cpTi and Ti-6Al-4V alloy.
140
140
4.39 Nyquist plots of the triple layer coated cpTi and
Ti-6Al-
4V alloy
4.40 Randle circuit forHAp tripl layer coated cpTi and Ti-
6Al-4Valloy
141
XIV
Al2O3 Alumina
EDS Energy Dispersive X-ray Spectroscopy
EIS Electrochemical Impedance Spectroscopy
FTIR Furrier Transformation Infra-Red
LSV Linear Sweep Voltammetry
OCP Octa Calcium Phosphate
XRD X-ray Diffraction
alloys by RF Magnetron Sputtering at
Different Thin film TiO2 Nano
Thickness
Magnetron - Sputtered Process And
Differentiate Between Single And
2015
1
1.1 Introduction
The natural synovial joints (hip, shoulder, knee or dental) are
performing
well under a wide variety of conditions. Unfortunately, these
joints are also
susceptible to inflammation and degenerative diseases that may
result in
pain and stiffness. Modern history of the world has huge
military
treatments including progress in old as well as wars that produce a
large
number of people s need replacing failed bones.
The typical cause of the degeneration is degradation in the
joint’s
mechanical properties that results from a failure of the normal
biological
repair processes, or excessive loading conditions; and hence, the
natural
joints can no longer function, they can be replaced by artificial
biomaterials
in a procedure known as a total joint replacement (TJR)
arthroplasty [1].
The studies in the biomaterial field are focused on Two
areas:
- The development of the new metallic materials with
mechanical
properties close to the human bone and
- The surface modification techniques enabling optimum
biocompatibility
(which is closely related to their corrosion behavior in
biological
environments) and Osseo integration of medical alloy
implants[2].
Osseointegration is defined as “the apparent direct attachment
or
connection of osseous tissue to an inert, alloplastic material
without
intervening connective tissue”[3].
The properties of alloys that used as orthopedic biomaterials are
high
specific strength and corrosion resistance protected from
accelerated
corrosion rate by natural passivation oxide layer created on the
surface and
Chapter One Introduction and Literature Review
2
acts like an electrical resistor to retard the anodic dissolution
of metal
cations. Therefore, all implant alloys have a finite, albeit slow,
uniform
corrosion rate in vivo. The damage occurs to this passivation
layer, such as
by fretting or wear, maybe produce conditions conducive to
accelerated
focal corrosion and failure[4].
Today, many different metal alloys are used in implant materials,
including
titanium alloys such as commercially pure titanium CPTi,
Ti-6Al-4V,
stainless steels, Co-Cr-Mo alloys,…etc. The use of these materials
dates
back to the middle of the 20 th century and use of titanium alloys
for
surgical implants dates are back to the 1940s[5].
The modulus elasticity for the cobalt-based alloys are higher than
that of
other materials such as stainless steels, while Ti-alloys have a
modulus( 96
GPa for cp Ti and 117 GPa for Ti-6Al-4V )and density about half
that of
the Co-alloys. The commercially pure titanium (cpTi) used in
implants
because it has an excellent tissue compatibility and better
corrosion
resistance than stainless steel, however, limited strength and very
poor
wear resistance hindered the material[6].
In the late 1970s, the Ti-6Al-4V alloy was used in prostheses grew
because
of its high excellent corrosion resistance due to an oxide film
grows
spontaneously on the surface upon exposure to air (repassivation
ability)
and good mechanical properties. Biocompatibility studies indicated
that the
chemical constitution of Ti-6Al-4Valloy is not perfect due to the
presence
elements of Al and V, undergoes rapid corrosion and ion that could
be
released in the human body, causing toxic effects environment leads
to
oxygen depletion and high concentrations of the element in the
surrounding
tissues. The vanadium generates cytotoxic reactions and aluminum
results
in the bones softening and neuron damage. The surface
modification
techniques were used to prevent the ion exchange in the body
environment
Chapter One Introduction and Literature Review
3
and enhance the properties of the alloys surface, corrosion
resistance and
passivation ability [7,8].
Hydroxyapatite, HAp, Ca10 (PO4)6(OH) 2, as a naturally occurring
mineral
found in the inorganic component of human bone is ahighly
bio-affinity
that can make or form the bio-ceramic coatings and is bonded
directly with
bones and that are required for Osseo integration of the
implant[9].
The structure, thickness and number of a HAp layers on surfaces
effect on
the quality of implants, coating with a single HAp layer, have
poor
mechanical, but much better bio-medical properties. Double and
triple
layers introduce a large number of interfaces, parallel to the
substrate
surface and may deflect micro cracks and wide different atomic
intermixed
between the substrate and the top-bio coating which improves of
the
adhesion strength.
HAp nanoparticles are one of the highest values in the biomedical
field, are
the major and most abundant material in human bones and teeth. As
a
result, synthetic HAp nanoparticles that mimic natural HAp are
extensively
synthesized to repair and substitute human bones [10].
Biocompatibility carried out by immersion test was performed on
coated
sample by immersed it in Simulated Body Fluid (SBF) for one
month.
Using this solution, it is possible to form biomimetic apatite on
bioactive
surfaces[11].
In vitro carried out by corrosion test, the resout of corrosion
behavior
implantspacimen when coated surface with different thickness in
SBF
solution, compared tothe corrosion behavior of uncoated implant
specimen.
The corrosion resistance can be evaluated by electrochemical
techniques,
namely:
1-OCP -Open Circuit Potential consists in a period during which no
current
can flow and no potential can be applied to the working
electrode
Chapter One Introduction and Literature Review
4
state conditions and
3-EIS-Electrochemical Impedance Spectroscopy.
Corrosion is a natural process related to the nature of the passive
film and it
is surface phenomenon, that converts refined metal to their more
stable
oxide formed on the material in the intended environment [7].
1.2 Literature Review
Ceramic have played an important vital role to improve Osseo
integration
properties of the implant. It can be divided into two main types,
oxides
ceramic (like ZrO2, Al2O3, ,TiO2,MgO ..etc.) and non-oxide
ceramics
(HAp, SiC, ZnS, Si3N4 .. etc.)[12] .
A-Coating with Titanium Oxide TiO2
Previous studies indicated the growth of apatite on TiO2 coated
specimens
after imation in SBF. It is possible to increase the range of
biomaterial
applications and reduce from the release ion surface by depositing
thin
biocompatibility layer of TiO2 ceramics on titanium alloys.
A.Dakka et al (1999) [13]
Have been studied the optical response (as a
function of oxygen partial pressure, target state and sputtering
power) of
TiO2 layers prepared by RF sputtering used substrate mature SiO2,
Si and
NaCl. They found that "TiO2 layers deposited from a used target
exhibit a
high absorptance" which decreases highly when introducing oxygen
partial
pressure, whereas increasing the sputtering power will lead to an
absorbent
TiO2 matrix.
has studied the effect of ceramic Thick coatings
TiO2, Al2O3 ceramics were deposited on NiTi alloys with high
power
lasers type CO2 by use powder injection to the melt pool using
laser
Chapter One Introduction and Literature Review
5
typical coating material with laser beam types and principal
characteristics
were reviewed. It was suggested that a better control of the
process and
better understanding of the physical phenomena occurring in
laser
treatment, such as injection of solid particles in a melt pool or
solidification
of the coating.
studied the deposited TiO2 thin film on Si and glass
substrates using RF magntron sputtering with the mixing of two gas
types
Ar and O2. They found that crystallinity and microstructure of thin
film
depend upon the depostion time and O2/(Ar+O2) mixing ratio.
The
crystalline phase occurs at high O2/(Ar+O2) ratio and long
deposition time;
while low mixing ratio amorphous phase occurred. The thickness of
thin
films increases with depostion time.
Wenjie et al(2002) [16]
Found that it anatase, rutile or amorphous TiO2 films
can be produced using RF magnetron sputtering deposited on the
glass
substrate with virus crystalline structures by varying charectrize
sputtering
such as gases, substrate temperature and annealing process.
Lei Miao et at(2002) [17]
Studied the polycrystalline and epitaxial TiO2 film
formed on silicon and pyrex substrates using RF magntron sputtering
with
TiO2 target in Ar and O2 gas. They conformed that formation of the
single
phase TiO2 can be achieved by changing the deposition parameters.
A
single phase rutil appears at fixed Ar pressure of 0.1Pa and 600C
heat
substrate, also appears at 325C while changed pressure from 0.1 to
1 Pa.
Whereas anataes phase appears at 1Pa, with Oxygen flow ratio 0.1 to
0.7
sccm.
Have prepared coating on Ti and Ti-6Al-4V
alloys substrate by Al2O3, HAp and TiO2 Powder using plasma
spraying
Chapter One Introduction and Literature Review
6
method. The results shows that "on the surface", the large
particles
compose increases porosity at increasing thickness of thin film
coated.
Xiaobing Zhao(2006) [19]
substrate using plasma spraying method. A chemical treatment method
was
used to increase surface bioactivity, the immersed sprayed TiO2
coating in
10M of NaOH solution at 60C o for 24h, then soaked in SBF to
create
bioactivity. A containing hydroxyapatite can be induced to form on
the
surface during the immersion in SBF in vitro test.
Maria C. Advincula et al(2006) [20]
investigated the properties of surface
titanium alloy properties by surface sol-gel processing (SSP) or
by
passivation with nitric acid to enhance the bio reactivity of
Ti-6Al-4V
substrate. Surface analysis of sol–gel–create oxide on Ti-6Al-4V
substrates
showed a prevalent titanium dioxide (TiO2) composition with
abundant
hydroxyl groups. The surface was highly wet table, rougher, more
porous
and cells adhered to the sol–gel-coated surface compared to that of
the
passivated substrate.
Coated the surface of Ti by a robust (lum thick)
mesoporous titania layer (MTL), and they found that the surfaces
can very
efficiently. These coatings are produced by anodization of Ti at
elevated
temperature in a glycerol/K2HPO4 electrolyte, followed by an
appropriate
etching process. Immersion tests in two types of simulated body
fluids
(Kokubo SBF and Bohner and Lemaitre SBF) are combined to
examine
these layers with regard to their ability to form hydroxyapatite.
The MTL
layers lead to a significant enhancement of HAp formation and
anchoring
in the structure compared with non-coated.
Chaing-Hua Wei et al(2011) [22]
studied the effect of surface roughness on
the crysralline strucur of TiO2 by deposited polycrystalline TiO2
thin films
on various unheated glass and ITO rough substrates using RF
magnetron
Chapter One Introduction and Literature Review
7
sputtering. The XRD results show that the polycrysralline structure
was
formed on the unheated glass while nano crystalline structure
easily
formed on the rough surface.
Krishna et al (2011) [23]
studied the effect of titanium oxide under layer
formed by DC sputtering on the crystalline phase of the TiO2 thin
films
formed on 316L stainless steel by non-reactive RF magnetron
sputtering of
TiO2 target. They used Ar gas of 0.5 psi and substrate temperature
of
300C, as deposition condtions. The resulting underlayer TiO2 films
have
tribological properties, wear and corrosion resistance of the
stainless steel
in terms of of the reduced of friction coefficient and enhancement
in the
surface hardness.
develop a bioactive and corrosion resistant by
converting the bio inert surface of NiTi to bioactive and
biocompatible
surface by depositing TiO2 particles on the NiTi surface
using
electrophoretic deposition process. TiO2 particles were prepared
using a
mixture of acetone and n-butanol (0%, 30%, 60%, 80% and 100%
acetone)
without using any dispersant.Surface morphology of coatings
shows
deposition within 0% acetone cause to crack-free and dense coating
with
relatively coarse grains and high corrosion resistance.
B-Coating with hydroxyapatite HAp
Coating with HAp which can reinforce rapid bonding to natural
bone,
there are different techniques that commonly uses HAp to coat
implant
material include: immersion-coating, dip-coating,
electrophoretic
deposition, hot isotactic pressing, solution deposition,
flame-spraying and
sputter-coating. The biological, chemical and mechanical properties
of the
coating depend on the method used. The immersion coating and
dip-
coating methods disadvantage may degrade compromise the purity of
the
hydroxyapatite and they require high temperature for the
post-sintering of
Chapter One Introduction and Literature Review
8
agglomerates& poor biological fixation to the metal substrate.
Hot isostatic
pressing is two-stage technique which is difficult to seal borders
on
implants and cumbersome with complex shapes. Plasma spraying
process
with high velocity and temperature technique produce high degree
of
melting to ceramic powder. Limitations associated with the
plasma-
spraying process included; high porosity, poor adherence and
cost[25].
Most sputter-coating techniques are slow and have an inadequate
rate of
deposit on a substrate that has been placed in a vacuum chamber,
leading to
overcome the contaminationencounter with other technigues.
Sputtering processes are common method to deposited thin film
properties
in ways not viable through other deposition techniques. This
flexibility
comes from many possible combinations and permutations of
deposition
process parameters, gas, target and sputtering equipment[26].
RF sputtering can be performed on all types of materials. One
primary
benefit of RF is the ability to deposit high melting point oxides
at relatively
low substrate temperature. RF sputtering is widely used in a
variety of
applications ranging from semiconductor industries to aerospace
[27].
H.M.Kim et al (1996) [28]
have prepared a sodium titanate layer on
commercially pure titanium and different titanium alloys substrate
using
simple chemical method.This method includes the socking of cpTi,
Ti-
6Al-4V, Ti-6Al-2Nb-Ta, and Ti-15Mo-5Zr-3Al alloys substrates in
10M
NaOH aqueous solution and then heat-treated at 600°C. The
treatment
resulted in the formation of a dense and uniform bone like apatite
layer on
surfaces after immersion in Simulated Body Fluid(SBF) with
ion
concentrations nearly equal to those of human blood plasma. The
alkali (in
NaOH) - and heat-treated (at 600°C) metals bond to living bone
through
the bonelike apatite layer formed on their surfaces in the
body.
Chapter One Introduction and Literature Review
9
used atmospheric plasma spraying to creat a
mixture of crystalline and amorphous phase of HAp on implants
to
enhance adhesion and fixation of the implant to the surrounding
bone.
Inthier study,heat treatment was used to convert the amorphous
phase to
crystalline followed by the diffusion of hydroxyl ions, the
amorphous phase
contracts, because of hydroxylation and crystallization, which
produce
cracks in the coating. They suggested that "the increasing in
crystallanityof
the coatings on the implant is not recommended"
J.G.C.Wolke et al(1998) [30]
RF magnetron sputtering.A copper disc provided with plasma sprayed
HA
was used as target material. The substrate was coated with a film
of
different thicknesses (0.1, 1 and 4µm) and annealed at 500C for 2h
.The
analysis results show that an amorphous film with low thickness 0.1
µm
changed in crystalline apatite, while amorphous film with high
film
thickness( 1 and 4 µm )sputter coating into amorphous–crstallinie
structure
shows some cracks. All the heat treated films showed stability
when
immersed in SBF. The authors concluded than the In vitro test
of
amorphous surface with low thickness dissolved after 4 weeks in
SBF
while high thickness only shows sign of surface dissolution.
Masayuki KON et al(2002) [31]
investigated a HAp ceramic surface
modification of the compositional gradient layer containing α-TCP
using
two-step treatment incloudes the immersion in5.0 mol/L H3PO4
solution
and buffered solution (pH equile 4.0) consisting of phosphate and
then
heated at 1250C for 1h. XPS analysis shows that HAp ceramic
increases
with increasing depth from the surface. The α -TCP transforms to
HAp in
water, in vitro or in vivo. Hence the surfaces of specimens can be
modified
with a compound such as α -TCP. However, the surface modified layer
or
compositional gradient layer is remarkably thin, with maximal
thickness of
Chapter One Introduction and Literature Review
11
approximately 2μm. It appears that the compositional gradient
layer
containing α -TCP on the surface of HAp has more effective
bioactivity
than the non-treated HAP ceramic.
Takayoshi Nakano et al(2002) [32]
studied the effect of heat treatment on
improvement of the microstructure and solubility of HAp (in vivo
test).
The change in phase layer of surface and control grain size of the
HAp
were exaimend to improve the the solubility of HAp .They found that
high
solubility and boundary layer composed of (TCP,αTCP,CaO) was
formed
by annealing HAp at high temperature 1350C in vacuum.
Ricardo M. Souto et al(2003) [33]
investigated the degradation
characteristics of hydroxyapatite (HA) coated on orthopaedic
Ti–6Al–4V
alloy using plasma spray .They studied the degradation by immersed
in
Ringer’s salt solution and using electrochemical impedance
spectroscopy(EIS). "The characteristic feature that describes
the
electrochemical behavior of the coated material by plasma spray is
the
coexistence of large areas of the coating itself with pores where
the
substrate is exposed to the aggressive media that is, the corrosion
resistance
of the biomaterial is not greatly affected by the presence of the
ceramic
coating, but rather depends on the passivation ability of the
metallic
substrate and, to a minor extent, on the porosity of the ceramic
coating".
Shuyan Xu et al(2005) [34]
studied the effect of time deposition tD on
thickness of thin film HA-coated on Ti-6Al-4Vorthopedic alloy using
RF
sputtering . The XRD intensity amplitude of HA crystalline
content
increases with increasing tD and becomes more pronounced.
Chun-Cheng Chen et al(2006) [35]
Studied the effect of heat treatment at
range of 500–700C on the improvement of bioactive
hydroxyapatite
(HAp)-coated Ti-6Al-4V alloy implants using plasma spray.
Hanks’
balanced salt solution was used in vitro corrosion test for samples
plasma
Chapter One Introduction and Literature Review
11
spraed with heated and non-heated. They concluded that "the
heat
treatment at 600C for 1h in air, endowing with increased
recrystallization
of amorphous calcium phosphate of as-sprayed HA coatings and
reduced
the defects without significantly reduced bond strength, provided a
better
corrosion protection than the other two treatment temperatures 500
and
700C".
studied CaP thin film(0.09-2.7μm) deposited
on pure titanium, 316L stainless steeland Ti-6Al-4V alloy by using
RF
sputtering. Ca-P coating prepared by RF-magnetron deposition has
high
values of adhesion strength, morphology increasing in roughness
with
increasing thickness of thin film without pores and micro cracks.
Also at
increasing thickness, the Ca concentration increases more than
P
concentration probable that during the sputtering the atoms of
calcium Ca,
which have a little more mass, displace the atoms of less mass, in
particular
phosphorus P. The cracks and damage of coated accure at thickness
more
than 1.6 μm.
worked on graded Ca-P bioceramic on a Ti-
6Al-4V orthopedic alloy by RF sputtering. They conclode that
chemical
composition and presence HAp, CaTiO3, and CaO mineral phases can
be
effectively controlled by the process parameters. For the film
synthesized at
700 W RF power, under 12 torr working gas pressure giving a Ca/P
ratio
of 1.64.Optical emission spectroscopy suggests that CaO + is the
dominant
species that responds to negative DC bias and controls calcium
content.
Biocompatibility tests in SBF at 36.5°C and a pH value of 7.4
confirm a
positive biomimetic response evidenced by soaking for 24h, some
small
apatite particles were formed indicating the apatite nucleation
onset and
for 48h, it was observed that apatite had covered completely the
surface.
Chapter One Introduction and Literature Review
12
investigated the physico/chemically
characterize of (Ca–P) thin film of 400–700nm thickness coated by
a
sputtering process on (Ti-6Al-4V) alloy. The corrosion resistance
for
coated and uncoated samples were determined and tested in
Phosphate
Buffered Saline (PBS). They found that by RF sputtered can produce
thin
bioactive ceramic with higher corrosion resistance compared to
uncoated
surfaces.
implant cpTi alloy by Hydroxyapatite (HAp), fluorhydroxyapatite
(FHA)
and fluorapatite (FHAp). Calcium nitrate and triethyl phosphite
were used
as precursors under an ethanol-water based solution. Increasing the
coating
speed decreases the porosity and thickness of the coatings,
dissolution
( and
rates) rates for ball rates could be decreased by increasing
the sintering temperature.
used the pulsed high power action on
electrode electrolyte interface with microplasma discharges method
for
obtaining Hydroxyapatite (HAp) and HAp/yttria-stabilized
zirconium
(YSZ) nanostructural bioceramic coatings on titanium and titanium
based
alloys for stomatology and orthopedy. Such coatings can be
uniformly and
smooth.
evaluated the effect of
temperature on the charastristic and isolation of HAp from tuna
bone at
different temperatures ranging from 200 – 1200°C. Based on their
analysis,
the formation of nanostructured HAp (80–300 nm) with beast
result
characterization at 600 °C and crystal agglomeration was observed
with an
increase in temperature due to the appearance of tricalcium
phosphate, the
isolation temperature between 600–900°C has tremendous impact on
the
production of HAp. Low temperature needed to formation HAp.
Chapter One Introduction and Literature Review
13
investigate the nature of apatite coatings on
implant Ti alloy prepared by the biomimetic method. The method is
based
on the soaking of solution PBS based process where the compositions
of
the soaking medium and thus the formed coatings can be controlled.
The
ions of solution exchange with ions of oxide layer; biologically
active ions
were substituted into the apatite coatings. The in vivo effects
showed that
the ion substituted apatite coatings have good biocompatibility and
can
promote early bone formation.
Rajan Anand (2012) [42]
sintered (900 for 2h) and non-sintered hydroxyapatite layer coated
on
Ti6Al4V alloy by using electrophoretic deposition carried out at a
constant
voltage of 30V for 5, 10 and 15 minute duration at different pH
values of
1.5, 2.5 and 3.5 respectively. SEM images show that at constant
voltage of
30V and low pH of 1.5, coated surface shows no cracks. XRD
analysis
showed that "before sintering, few other compounds were present in
HAp
powder while after sintering they were not oxidized.The more
coating
layers, leads to a wide different atomic intermixed between the
substrate
and the top-bio coating which improves the adhesion strength.
C-Coated with Different doublelayer
studied the behaviour of Ca/P enhanced by
MgAl2O4 in (SBF) with PH value 11-12. The results show that at
low
concentration of MgAl2O4, the formation of apatite increases on the
surface
of Ca/P. The formation composite ceramic in SBF depends on
dissolation
of the surface Ca/P and amorphous material produces from reaction
of HA
with Mg Al2O4.
worked on the development of biocompatiblity
and improved the interface properties between the coating and the
substrate
coating by multi-layered concept. Ti and HAp layers deposited on
Ti-6Al-
Chapter One Introduction and Literature Review
14
4V substrate by using RF magnetron sputtering, an underlying Ti
bond
coat, the alternating layer, and an HA top-layer. For corrosion
test,
itpresented multi-layered coatings exhibited a better immersion in
SBF
than single layer HA coatings and high adhesion on substrate.
Yu-PengLu et al(2003) [45]
used two layer hydroxyapatite
(HAp)/HAp+TiO2 to coat titanium . Three kinds of specimens,pure
HA
coating, 50 vol% HA+50 vol% TiO2 composite coating and two-layer
HA
top coating/ bond coat composite coating on titanium were produced
by
plasma spraying followed by heat treatment at 650C for 120 min
which
resulted in transfer calcium phosphates into HA. From EDS, there
exists
inter diffusion of the elements but from XRD no chemical
product
between HA and TiO2, such as CaTiO3 was formed. The positive effect
of
the composite coating is the decrease of residual stress in HA top
coating.
The Ti substrate bonds well via its TiO2 hobnobbing with the Ti
oxides
formed on Ti substrate.
R.Tomaszek et al(2004) [46]
investigated the effect of changing the mixing
ratio of theTiO2 and Al2O3 on microstructure of coated layer. The
results
showed that the mixing Al2O3 with 13%TiO2 give the best ratio
microstructure for coating titnium alloys by using plasma
spraying.
Zuqiang Qi (2004) [47]
Al/Al2O3,Ti/TiN have high potential for nourmas applications .
These can
exhibit enhanced properties due the effects on mechanical behavior
of the
multilayer. The Al in the multilayer has finer grains and the
interface
between Al and Al2O3 layers is continuous without any major defects
than
those when deposited as single layer, Ti layer grows in coarse
columns
when deposited on TiN with sharp intephase between metal and
ceramic.
Hae-Won Kim et al(2004) [48]
developed (HAp/TiO2) coated on to Ti
substrate by a sol–gel method with thicknesses of approximately 800
and
Chapter One Introduction and Literature Review
15
200nm. The HA coated up layer to enhance the bioactivity and
osteoconductivity of the Ti substrate and, TiO2 layer was inserted
mid to
improve the bonding strength between the HA layer and Ti substrate
as
well as to increasing corrosion resistance of the Ti substrate,
With heat
treatment thin film at 500 C. In vitro test, osteoblast-like HOS
cells grew
on HA/TiO2 in a similar fashion but a little more actively compared
to
those on the bare Ti and TiO2 coated Ti. In vivo corrosion test,
with a
physiological saline solution at 37C, the TiO2 coated improved
corrosion
resistance of the Ti substrate.
Chun-Cheng Chen et al (2005) [49]
worked on deposit HAp as outer layer
and TiO2 under layer on Ti-6Al-4V substrate using plasma spraying
to
improve the coating–substrate interface properties. The results
show that
graded (HA/TiO2) had better mechanical properties than single
HA
coatings, surface chemistry and morphology of the graded coatings
were
similar to those of single layer HA coatings. The in vitro
electrochemical
measurement results after immersion in HBSS for 1h,also indicated
that the
graded coatings seem to possess a better corrosion-resistant
ability behavior
than single HA coatings.
carried out electrophoretic co-
deposition of HAp and Al powder to form HAp/Al2O3 composite
coating
on cpTi alloy. The composite coating sintered at 850 ,and due to
the
thermal decomposition, the HAp phase appeared with no cracks and
the
bonding strength to the substrate was significantly improved
compared to
the single HAp coating.
Wisanu Boonrawd et al(2010) [51]
deposited composite of HAp/Ti target
on silicon wafer and Ti-6Al-4V alloy substrates by using RF
magnetron
sputtering. The composite of HAp/Ti films annealed at 600°C and
700°C.
The results showed that the filme exhibited high adhesion strength
as
Chapter One Introduction and Literature Review
16
compared to monolithic HAp coatings on Ti-base substrates and the
XRD
analysis revealed peaks of HAp and calcium titanate (CaTiO3).
I. Mercionu et al(2012 )[52]
obtained new HA/Y2O3:αAl2O3 system by
deposition of a hydroxyapatite thin layer with thickness of (~150
nm),
using RF magnetron sputtering, onto sintered substrates of low
yttria (150
ppm) doped alumina. The morphology of the layers was
homogeneous,
while at the HA/Y2O3:αAl2O3 showed a strong structural
interdependence
between the sintered substrate interface and covering layer with
no
loosening phenomena were evidenced. The grain size showed a
linear
increase with the annealing time.
L. Mohan et al(2012) [53]
studied the corrosion and scratch behavior of
TiO2 + 50%HAp nano ceramic coated on Ti–13Nb–13Zr orthopedic
implant alloy using (EPD) sintering at 850 C. The corrosion
behavior of
the coatings was evaluated using SBF-Hank’s solution. The
sintered
coating exhibited higher density, adhesion and lower porosity
compared to
unsintered samples, and higher corrosion resistance compared to
the
substrate.
Conclusion Remark of Literature Review
1-Heat treatment was use to convert the amorphous phase to
crystalline
followed by the diffusion of hydroxylions .By increasing time of
heat
treatment wich lead to produce crack also formed layer
composed
(TCP,CaO).The cracks and damage of coated CaP accuer at thickness
more
than 1.6um .
2-The composite (HA/Ti) films annealed at 600C ,the film exhibited
high
adhesion strength with composit (CaTiO3).The HA/YeO3:αAl2O3
system
deposited by using RFsputtering ,showed a strong structural
Chapter One Introduction and Literature Review
17
biocompatibility layer.
3-Immersion tests for coated samples in (SBF)are combined to
examine the
coated layers with regard to their ability to form HAp.
4-Any change in RF sputtering conditions working effected on
the
characterize of thin film.
Present work will
Deposite TiO2 and HAp as single layers ,TiO2+ Al2O3 as multilayer
and
TiO2+ Al2O3+HAp as three layer on cpTi and Ti6Al4V alloy
using
RFmagntron sputtering .Structure analysis and corrosion tests will
be used
to evaluate the properties of single and multilayers also the
coating will be
achieve biomimetic test to see the efficiency of the
compatibility.
1.3 Thesis Aims
1-Develop biocompatibility coating on the surface of cpTi and
Ti6Al4V
alloy using ceramic nanoparticle TiO2,HAp and Al2O3 by using
RFmagntron sputtering processing .
2- Enhance the interface properties between the coating and
the
substrate by using multilayer of different bioceramics .
3-Evaluation the corrosion resistance.
Chapter Two Theoretical Aspects
CHAPTER TWO
Theoretical Aspects
2.1 Introduction
This chapter provides a review of selected biomaterial alloy,
methods of
modification surface, coating by ceramic using modern technique and
basic
corrosion process. It starts with an introduction to the most
impotent
properties of titanium alloy as implant material, important of
modification
surface using ceramic as HAp, Al2O3, TiO2 with general properties
as
biocompatibility coating, the advantage of using plasma RF
sputtering
technique for enhance the properties of surface and corrosion
resistance
followed by basic corrosion process and corrosion kinetics,
descriptions
of some major electrochemical techniques for corrosion rate
evaluations
such as(OCP, LSV, and EIS).
2.2 Biocompatibility Requirements for Joint Replacement
Materials
Biomaterial is defined by the interaction between the body and
the
material; the effect of the material on the body and the effect of
the body
environment on the material [54]. The advent of an aseptic
surgical
technique developed by Dr. J.Lister in the 1860s mad the
biomaterials
become practical. The Consensus Conference of the European Society
for
Biomaterials (1986) defined biomaterials as (non-living) materials
used in a
medical device, intended to interact with biological systems; they
have
many characteristics including mechanical, physical, chemical
and
biological properties that make it suitable for safe, effective,
and reliable
use within a physiologic environment, an environment that is
both
extremely feudal and yet sensitive to unforgiving irritating
foreign bodies
also, the materials must be able to perform their functions within
the body
Chapter Two Theoretical Aspects
81
environment for a long period of time (over 25 years) without
decay. Bone
is a living anisotropic and viscoelastic material that adjust to
the loads
placed on it, the elastic modulus of the material is the one of the
most
critical mechanical properties in implant applications with a low
value that
approaches that of natural bone which is critical in order to
provide good
load transfer between the implant and surrounding bone to spur
new
growth. Metal ions from metal implants are entry into the solution
of body
through an open wound and its follow corrosion by interaction with
the
surrounding tissues through reduction-oxidation (redox) reactions.
The
products of these reactions and the kinetics are determining
factors that
called the biocompatibility of an element. The titanium alloys are
highly
desirable for implant applications, high biocompatibility and
inertness of
titanium is due to the characteristics of its oxide layer. The
oxide layer in
titanium alloys represent in TiO2 layer is an excellent insulator,
and prevent
transmit electrons from elements alloy to body tissues and fluids.
Also,
because the oxide layer forms by oxygen diffusing inward towards
the
metal interface instead of metal ions diffusing outward and
combining with
oxygen there is limited accumulation of ions into the surrounding
tissue.
Ni, Cu, Au, Mo, Ti, Nb, Ta, Zr, and Pt have the capacity to be used
as
alloying elements for improved biomedical titanium alloys, Fig.
2.1
describing the biocompatibility of some of the various elements
[5].
Polarization resistance is the ratio of the applied potential and
the resulting
current response. This "resistance" is inversely related to the
uniform
corrosion rate, the higher the polarization resistance, the lower
the
corrosion rate [55].
biocompatibility of various metals [3].
2.3 The Basic Properties of Titanium
Titanium is one of the transition elements in group IV and period 4
of
Mendeleef’s periodic table. It has an atomic weight of 47.9 and
atomic
number of 22. Being a transition element, titanium has an
incompletely
filled d shell in its electronic structure. The incomplete shell
enables
titanium to form solid solutions with most substitution elements
having a
size factor within ±20% and in the element form, el ing in
1668 C).The properties of titanium alloys are principally by two
factors:
the chemical composition and microstructure. The chemical
composition of
the titanium alloys primarily determines the properties and volume
fraction
of the phases, hexagonal closely packed crystal structure (hcp) and
β
body centered cubic structure (bcc). Titanium alloys falls into
three classes:
α-all ys, α+β all ys and β-alloys. a e e a e f 882 5 C,
titanium
transforms from to β. Titanium alloys may be classified as
neutral,
Chapter Two Theoretical Aspects
08
and β depending up on the room temperature microstructure. The
main
properties that make titanium attractive for a variety of
applications:
1-Mechanical properties: Titanium is used in orthopedics due to
high
specific strength and low elastic modulus, wear, density and
abrasion
resistance because of its low hardness, this problem solved by
surface
modification to enhance the abrasive wear corrosion
resistance.
2-Biological properties: it is generally regarded to have
good
biocompatibility, relatively inert and have good corrosion
resistance
because of thin surface oxide when compared to more
conventional
stainless steel and cobalt-based alloys [5, 56].
2.3.1 Commercial Pure Titanium cpTi
F67 cpTi alloy have grades from 1to 4 are based on the low
temperature,
phase (hcp) the four different grades of cp Ti differ with respect
to their
oxygen content of cpTi affects its yield and fatigue strength. The
type of
grade is 2 the chemical composition (Ti, Fe, N, C and O),
98.9-99.6%
titanium concentration, 0.25 oxygen concentration and the yield
strength is
about 550MPa as shown in Appendix A.
cpTi has excellent corrosion resistance, fabric ability but poor
in
mechanical properties, especially strength. There are growing
numbers of
applications such as pressure vessels and dental implant. The
mechanical
properties are an equally important criterion for materials
selection [56].
2.3.2 Titanium Aluminum Vanadium Ti-6Al-4V alloy
F 136 Ti-6Al-4V for orthopedic load –bearing implants alloying,
addition
of 6% aluminum with phase and 4% vanadium with β phase to cpTi
as
shown in Appendix B, results in any alloy having good
mechanical
Chapter Two Theoretical Aspects
00
properties yield strength is about 880MPa (increasing the yield and
tensile
strength, acceptable elongations ) and chemical inertness. From
Fig.2.2 it is
clear that for this aluminum level of about 6% the / β,
transformation
temperature of 882°C for pure titanium is increase to about 1000°C
for the
two phase region + β. Since vanadium is a β -stabilizer, this means
that
the higher temperature β more readily transforms to during cooling
after
processing, and also during any subsequent heat treatment at
relatively low
temperatures, annealing at 700–750C.
The most popular titanium alloy, more than 50% of alloys in use
today are
of this composition [7,65].Ti- 6Al-4V alloy has been in clinical
use since
the 1950 there are reasons: good balance of its properties,
ultimate
strength, corrosion resistance, ductile behavior for loads
exceeding the
yield stress (thus reducing the risk of catastrophic brittle
failure), fabric
ability, available and cheap.
Figure 2.2: Phase diagram of Ti-Al-V alloy [57]
The passive film consists of TiO2 as the major component and
Aluminum
oxide (Al2O3), vanadium oxide (VO2) respectively on the metal
surface. In
return to VO2 is thermodynamically unstable and performs into
solution
body. An inert fibrous capsule formation around the implant to
prevent
Chapter Two Theoretical Aspects
02
damage to surrounding tissues and eliminated by body within 24
hours, this
is presumably why implants made of Ti-6Al-4V alloy have not shown
any
serious disadvantages so far [7] . In general, there is a problem
in using Ti-
6Al-4V alloy for implant applications belong to the large
modulus
mismatch between the Ti-6Al-4V alloy (~110GPa) and the bone
(~10-
50GPa), which could cause insufficient loading of the bone adjacent
to the
implant. Development of new alloy materials and modification of
the
surface of the currently used Ti alloys have been widely explored
to
overcome these problems. Either the development of Ti6Al7Nb
and
Ti5Al2.5Fe, where Nb and Fe were substituted for Vin Ti-6Al-4Valloy
as
less toxic alternatives or numerous bioactive surface modifications
of
titanium alloys for biomedical applications are very important for
achieving
further developed biocompatibility [58, 7] .
2.4 Surface Modification
The bulk properties of biomaterials, such as non-toxicity,
corrosion
resistance or controlled degradability, modulus of elasticity, and
fatigue
strength have long been to recognize to be highly relevant in terms
of the
selection of the right biomaterials for specific biomedical
application. The
events after implantation include the interactions between the
biological
environment and artificial material surface, onset of biological
reactions, as
well as the particular response paths chosen by the body.
In implants made of titanium, the various surface modification
techniques
required because:
1. In implants made of titanium, the normal manufacturing
steps
usually lead to an oxidized, contaminated surface layer that
is
surfaces are clearly not appropriate for biomedical applications
and
some surface treatment must be performed.
Chapter Two Theoretical Aspects
02
2. The specific surface properties that are different from those in
bulk
are often required. The difference between the modulus of
elasticity
of the metallic implant of Ti and the modulus of elasticity of
bone
tissue can be controlled with the fabrication of the structure
surface
of Ti implants.
4. Improve specific surface properties required by different
clinical
applications.
The surface modification techniques are a general concept that can
be
divided into surface treatments and surface coatings, or a
combination of
both, surface modification classified into [56]:
1. Mechanical methods: grinding, polishing, involve physical
treatment, shaping, or removal of the materials surface. The
typical
objective of mechanical modification is to obtain specific
surface
topographies and roughness, and to remove surface
contamination.
The deposition techniques that represents the processes to build
thin
film on substrate thin film have a thickness anywhere between a
few
nanometer to about 100 micrometer. These techniques can be
classified in two groups chemical and physics methods as shown
in
Figure 2.3.
respectively at the interface between titanium and a solution.
(CVD)
is a process involving chemical reactions between chemicals in
the
gas phase and the sample surface resulting in the deposition of
a
non-volatile compound on the substrate. In the sol–gel
process,
Chapter Two Theoretical Aspects
02
chemical reactions do not occur at the interface between the
sample
surface and solution or gel, but rather in the solution.
3. Physical methods: refer to such methods as thermal spraying
and
physical vapor deposition (PVD), where chemical reactions do
not
occur. In this case, films or coatings on titanium and its alloys
are
mainly attributed to the thermal, kinetic, and electrical energy.
In the
thermal spraying process, the coating materials are thermally
melted
Into liquid droplets and coated to the substrate at a high speed
(kinetic
energy). The generation of atoms, molecules or ions from targets
can be
accomplished by resistance heating, electron beam, sputtering and
laser
or electrical discharge in vacuum. Glow discharge plasma
treatment
and ion implantation are also categorized as physical methods [56,
59].
Chapter Two Theoretical Aspects
2.5 Plasma and Plasma Properties
Plasma can be defined as a gas containing charged and neutral
species
showing collective behavior. It consists of electrons, positive -
negative
ions, atoms, radicals and molecules. On average plasmas are
electrically
neutral, since any charge imbalance would create electric fields
that would
tend to move the charges to eliminate the imbalance. As a result,
the
combined density of electrons and negative ions will be equal to
that of
Chapter Two Theoretical Aspects
02
positively charged ions in any given volume of plasma. This
property of
plasmas is called quasi-neutrality [56].
Sputtering is a preparation technique where the ejection of a
material is
due to the transfer of energy from an energetic particle to a
surface. The
energetic particles, in the mixture form of ionized atoms or
clusters of
atoms, is passed to a substrate where a film of the target material
is
deposited. This process has been utilized to more conveniently
deposit a
wide range of materials since the momentum exchange is a
physical
process as opposed to a chemical or thermal process [56, 61].
The target is mounted opposite to the substrates in a vacuum
chamber. A
strong potential difference is applied in a gas, generally of
argon, with
possibly reactive gases (O2, N2, etc.), a precursor gas argon is
used in most
applications because of its mass compatibility with materials of
engineering
interest and its low cost. When the target is powered negatively,
typically
between 0.5 and 5 kV, ionized argon atoms provides the ion
bombardment
of the target.
The process is then, the ejection of the target atoms is resulted
from the
argon ion bombardment, their transfer to the substrate with a
kinetic energy
and the nucleation and growth of the thin film on the substrate
surfaces [3],
as shown in Fig.2.4.Magnetron sputtering (MS) and RF is a very
powerful
technique which is used in a wide range of applications due to its
excellent
control over thickness, preparing large area uniform films,
adherence of the
films. The magnetic field in the magnetron is oriented parallel to
the
cathode surface the local polarity of magnetic field is oriented
such that
E×B drift of the emitted secondary electrons forms a closed loop.
Due to
the increased confinement of the secondary electrons in this E×B
drift loop,
plasma density will be much hair compared to a DC or diode device.
The
Chapter Two Theoretical Aspects
01
result of the high plasma density and its proximity to the cathode
is a high
current, relatively low voltage discharge [62,59].
Plasmas have two important characteristics that are ideal for thin
film
processing:
Firstly, they are sources of chemically active species, radicals
and ions that
result from the collision of neutral background gases with
electrons which
are sufficiently energetic to break their chemical bonds. Substrate
surface
reactions can thus take place at much lower temperatures than in
thermal
processing.
Secondly, low-pressure plasmas offer the added advantage of a
non-
collisional sheath, ensuring that the energy of the ions striking
the surface
can be accurately controlled within a broad range. Energetic ions
play a
synergetic role in deposition and etching processes and a
determining role
in sputtering processes [63].
Chapter Two Theoretical Aspects
RF Sputtering Technique
DC methods cannot be used to sputter non-insolating targets because
of
charge accumulation at the target surface. The application of
high
frequency voltage can overcome the problem of charge accumulation,
this
solution by using RF methods typical voltage frequency is 13.6MHz
since
this has been allocated by the Federal Communications Commission
(USA)
for industrial-scientific-medical purposes. Sputtering
non-conducting
materials is based upon the fact that a self-bias voltage, negative
with
respect to the plasma potential, develops on any surface that is
capacitively
coupled to a glow discharge. The basis for this potential, which
forms as a
consequence of the difference in mobility between electrons and
ions, is
illustrated schematically in Fig. 2.5[64].
Figure 2.5: Schematic illustration of the development of a negative
bias in a RF
system : (a) The current-voltage characteristic for an electrode
immersed in
plasma,(b) When an alternating voltage is applied to an electrode,
a
positive/negative potential appears on the surface [64].
Chapter Two Theoretical Aspects
2.6 Ceramic
Ceramic materials use to cover the implants which used in
medical
applications, including bio inert ceramics (i.e. alumina and
zirconia), and
bioactive ceramics (i.e. calcium phosphates, titania, bioactive
glasses and
glass ceramics). Alumina excellent mechanical properties is for the
load-
bearing applications, while the bioactivity ceramics leads to the
potential
for Osteoconduction [65].
2.6.1 Hydroxyapatite HAp
Bone is formed of cells 10% and matrix 90% (organic 40% and
inorganic
60%). An organic matrix composed mainly of type1 collagen and
a
mineralized inorganic matrix (crystals of HAp and CaP). The
collagen
fibers which form bone are the result of the bonding by means of
crossed
links of a triple helix of chains of this material [66, 67], as
shown in
Fig. 6.2.
(a) (b)
Figure 2.6: (a) Bon Polypeptide triple helix tropocolagen bond
together forming,
(b) Most HAp fills in holes in collagen[66].
Chapter Two Theoretical Aspects
28
The mineral apatite, can be represented by the formula, M10(ZO4)6X2
(or
M5(ZO4)3X), as a general term that can be applied to the crystal
structures.
Each component (M, ZO4, and X) of the common equation
M10(ZO4)6X2
can be replaced by a large number of different elements or solid
states. M
or X can also be absent. The most common form found in nature is
calcium
phosphate apatite, whereby the M and ZO4 are Ca 2+
and PO4 3-
groups.
When the X is OH - , Ca10(PO4)6(OH)2, it is given the name
hydroxyapatite.
HAp is the key inorganic component of the hard tissues of
vertebrae, and is
an important substance in bioactive ceramic materials, it is one of
many
types of calcium phosphate (or calcium orthophosphate) as shown in
Table
2.1, but there are several others in this class. The simplest way
to arrange
these into order is by the calcium-phosphorus composition ratio.
At
physiological pH (7.2-7.6), HA is most stable.
OCP is an intermediate that appears during extraction of the HA
phase.
Wi hin TCP, α-TCP shows a stable phase between 1180 and 1430°C
and
ansf s in i s β- hase a highe e e a es α-TCP, a hypothermic
phase material, is often used in the same way as HA, in bioactive
ceramics.
Its absorptivity when implanted into the body is significantly
higher than
that of HAp. A a e ial c sed f a ix e f HA and α-TCP has also
been developed, called biphasic calcium phosphate (BPC).
A metastable amorphous phase material known as amorphous
calcium
phosphate (ACP) exhibits a significantly higher bioresorbable
property
than other crystalline calcium phosphate compounds [68, 69].
2.6.1.1 The Crystal Structures of HAP
A schematic diagram of the structure of HA as shows in Fig. 2.7,
the
structure shown here is the stoichiometric composition of HA where
by the
unit cell is composed of Ca10(PO4)6(OH)2 . The core framework of
the
Chapter Two Theoretical Aspects
tetrahedron, with two types of
channel structures that lie parallel to the c –axis at the
positions, (x, y) =
(0, 1/4) and (x, y) = (1/3, 1/12). The OH ion (z = 0 ± 0.2, 1/2 ±
0.2) and
Ca 2+
ion (z = 0, 1/2) lie within each of these structures. This Ca
2+
is
exposed on the crystal surface, thus playing a large role in the
physical
properties of HA, such as surface charge, and in interactions with
organic
compounds.
The stoichiometric crystal structure of HA has a monoclinic system
with
P21/b as its space group by Ca/P ratio of 1.67, exhibiting
lattice
parameters, a= 9.4215A, b= 2a, c= 6.8715A, gamma =120º.
Earlier
studies showed that it also could exhibit the hexagonal structure
with space
group P63/m. However, it is difficult to obtain the exact
stoichiometry
(Ca/P ratio) in HA because of different Ca/P ratios that can be
stabilized
depending on the synthesis method and conditions employed
[70,71].
Biological apatite is a non-stoichiometric form of HAp,
characterized by
Ca 2+
The trace elements include positively charged ions, Mg 2+
, Na + , K
the
ion which constitutes 5-8% of apatite in bone, by weight. The
ion can replace all of the OH - and certain
within the HAp
structure, termed A-type and B-type replacement, respectively.
B-type
replacement is common in bones, and is an essential factor in
altering the
melting point of the biological apatite. Increases in positive
charge with the
replacement of by
2+
Chapter Two Theoretical Aspects
22
Fig. 2.7 shows the crystal structure of hydroxyapatite, where the
c-axis is
perpendicular to the plane of the paper the x, y, z that are
mentioned in the
text correspond to the direction of a, b, c in the diagram with
defining the
repetitive cycle as 1. The numbers in this diagram are the
coordinates of the
z-axis. The hydrogen atom of OH - ions on the y = 3/4 plane are
hidden
behind the oxygen atom [39, 68].
Table 2.1 :The major calcium phosphates [68]
calcium phosphates material Chemical composite Ca/P
Monocalcium phosphate anhydrous(MCPA) Ca(HPO4)2 0.50
Monocalcium phosphate monohydrate
be a icalci h a ha e β-TCP) Ca3(PO4)2 1.50
Hydroxyapatite(HA) Ca10(PO4)6(OH)2 1.67
Chapter Two Theoretical Aspects
22
(a)
(b)
Figure 2.7:(a) and (b): The crystal structure of hydroxyapatite
hexagonal [39]
2.6.2 Titania TiO2
There are three phases of oxide titanium or titania in atmosphere
air are
anatase, rutile and brookite. Pure Rutile is desirable phase in
clinical
applications due to the high corrosion protection of metallic
implants, a
large single crystals can be easily obtained, high biocompatibility
and
usually stated to be the thermodynamically most stable form of
TiO2.
Anatase phase also has bioactivity properties. Anatase gradually
transforms
to rutile depending on temperature that has bioactivity properties.
There are
Chapter Two Theoretical Aspects
22
only small differences in Gibbs free energy between anatase,
brookite and
rutile (4-20 kJ/mol) meaning that the metastable polymer phase is
almost as
stable as rutile at normal pressures and temperatures. [14,
72].
2.6.2.1 The Crystal Structures of TiO2
The most important properties of TiO2 is that it has the crystal
structure
same as the crystal structure of titanium alloys; so, it is useful
to be used
for coating it. The structure of rutile and anatase can be
described in terms
of chains of TiO6 octahedral.
The two crystal structures differ by the distortion of each
octahedron and
by the assembly pattern of the octahedral chains. The unit cell of
rutile
contains two Ti atoms situated at (0, 0, 0) and (½a, ½a, ½c), and
four
oxygen atoms. The oxygen atoms form a distorted octahedron around
every
Ti cation.
The unit cells of anatase contains four Ti atoms located at (0, 0,
0), (½a,
½a, ½c), (0, ½a, ¼c), and (-½a, 0, -¼c), and eight oxygen atoms.
In
anatase, too, the Ti cation is surrounded by a distorted oxygen
octahedron,
the unit cell structures of the rutile and anatase crystal are
illustrated in
Fig.(2.8 a and b) respectively and the physical properties are
tabulated in
Table (2.2)[73,74].
22
Table: 2.2 Some of the physical properties of anatase and rutile
TiO2 structures
at room temperature[37]
u=0.3048 Å, c/a=0.6441
Tetragonal
u=0.2081 Å, c/a=2.5143
Energy gap Eg 3.3 eV 3.0 eV
Mass densi y ρ 3.894 g/cm 3
4.23 g/cm 3
(a) (b)
Figure 2.8: (a) rutile and (b) anatase crystallographic unit cells.
Small spheres
represent Ti atoms, large spheres represent oxygen atoms
[73].
2.6.3 Alumina Al2O3
It is a very hard material, wear resistance and its hardness is
exceeded only
by diamond, high melting point, i.e, above 2000°C (3632°F). Alumina
is
also an inert substance, and at room temperature, it is insoluble
in all
Chapter Two Theoretical Aspects
22
ordinary chemical reagents. These qualities make it useful as a
biomaterial.
For example, alumina is used for artificial joint replacements;
porous
al ina is sed as a ‘b ne s ace ’ and f ee h i lan s The s
important and wide-range use of alumina is in the field of ceramics
[75,76]
2.6.3.1 The Crystal Structures of Al2O3
The basic unit cell structure of corundum (the naturally Occurring
alumina)
is hexagonal. The forms of metastable Al2O3 s c es, α, ρ, γ, η, θ,
χ and
k Al2O3. Those kinds of transition Al2O3can be produced from
heat
treatment of aluminum salts. The c ys al s c e α- Al2O3 which is
called
corundum structure, ideally consists of close packed planes of
large oxygen
anions (radius 0.14nm) stacked in the sequence as shown in Figure(
2.9)
,the aluminum cations (radius 0.053 nm) have valence of +3 and
oxygen
anions have valence of -2.There can be only two Al 3+
ions for every three
Figure2.9: The crystal structure of α alumina [76]
2.7 Biomimetic and Porosity
Biomimetic are based on the nucleation and growth of calcium
phosphate
in simulated body fluid, to produce an apatite layer on the surface
of Ti
implants, increasing their osteoconductivity and consequently
favoring
Osseo integration. Bioactive coating on the porous surface is an
attractive
method to improve the quality of the bone-implant interface.
Porosity can
be defined as the percentage of void spaces in a solid. The
Osseo
integration obtained with porous Ti is achieved by bone growth into
the
pores, called "bone ingrowth", which improves micromechanical
interlocking, the interlacing of bone tissue within the implant,
preventing
mobility since it increases the contact area between the
biomaterials and
bone tissue, resulting in improved implant stability over time, as
well as
accelerating the process of Osseo integration.
The porous structure must be produced with high porosity to
provide
sufficient space for cell adhesion and subsequent formation of new
bone
that permits the transport of body fluids and the proliferation of
new
vasculature, while providing adequate mechanical properties to
withstand
stresses during implantation [60].
2.8 Corrosion
Corrosion is the most important example of surface chemistry. It
is
degradation and transformation of a metallic structure into other
chemical
structures due to the fact that the interaction with their
environment and
corrosion of most metals is inevitable [77].
Corrosion of metallic implants causes many problems like:
1. Failure of implant system belongs to additional ions into the
body
2. Corrosion fatigue causes weaken for implant.
Chapter Two Theoretical Aspects
3. Generation of particulate debris, which may aggravate the
body
environment.
Corrosion of metal implants is critical because it can negatively
affect the
mechanical properties and biocompatibility. The materials used must
be
stable retaining their functional properties and not cause
any
counteractive biological reaction in the body. Corrosion and
surface
oxide film dissolution are the two mechanisms which introduce
additional ions into the body. The first condition for a material
of any
type that is to be used in the body is corrosion because metal ion
release
takes place mainly due to corrosion of surgical implants. It is
desirable to
keep the metal ion release to a minimum by the use of
corrosion-resistant
materials.
1. The material of implant itself (e.g. the chemical
composition,
microstructure, surface condition),
3. The construction (e.g. presence of crevices).
Changes in these variables can have a further influence on the mode
and
rate of metal ion release [78].
2.8.1 Gibbs Free Energy
Gibbs free energy G determines the feasibility of a chemical
reaction. A
react can take place spontaneously when the difference of free
energy is
nega ive ΔG < 0. The change of Gibbs free energy per mole of a
reacting
species is proportional to the difference of electrical en ials ΔE
and
:
Chapter Two Theoretical Aspects
22
Each mole of a reacting species of valence z transports a charge
of
Q = z F.
The negative sign in equation (1) is used to conform to the
convention
that a positive potential results in a negative free-energy change
for a
spontaneous reaction. The more nega ive he val e f ΔG, he g ea e
he
tendency for the reaction to go [79].
2.8.2 Electrochemical Kinetics of Corrosion
Corrosion processes are controlled both by thermodynamics and
kinetics,
which are about the fundamental feasibility of the reaction and the
rate,
respectively. It is important to know how fast corrosion takes
place. So,
to develop more corrosion-resistant alloys and to improve methods
of
protection against corrosion, an understanding of the fundamental
laws of
electrochemical reaction kinetics is essential.
The rate of electron flow to or from a reacting interface is a
measure of
reaction rate. The proportionality between current, I, in amperes,
and
ass eac ed, , in an elec che ical eac i n is given by Fa
aday’s
Law [78]:
m
(2)
Where is the atomic weight (11.97 amu for titanium), and t is the
time.
Dividing equation (2) through by t and the surface area, A, yields
the
corrosion rate;
Where j, defined as current density (I/A). This equation
shows
proportionality between mass loss per unit area per unit time
(mg/cm 2 /day) and current density (µA/cm
2 ). Units of penetration per unit
Chapter Two Theoretical Aspects
28
time result from dividing equation (3) by the density, D(4.5 g/cm 3
for
cpTi and 4.4 g/cm 3 for Ti-6Al-4V) of the alloys,(z) is the number
of
equivalents exchanged (1 for titanium) . For corrosion rate in
millimeter
per year, equation (3) becomes
) (4)
The biocompatibility of a material interaction with surrounding
medium
under real conditions of exposure as well as the resulting
deterioratiotion
of both components. Alloys proposed for implants systems can
be
evaluated using in vitro test corrosion direct current (DC) and
alternating
current (AC) or impedance electrochemical test methods.
2.9.1 Open Circuit Potential (OCP)
It is one of the simplest electrochemical tests available,
measuring the
open circuit potential (OCP), also referred to as the equilibrium
potential
or corrosion potential. The OCP is by definition the electrical
potential
difference between two conductors in specific electrolyte with
zero
current flow between them [81]. Monitoring OCP over time can
provide
vital information about the system being studied. And the system
has
reached a steady state and when transitions between different
states, such
as a passive and trans passive behavior will occur. Properties of
the oxide
formed on the test electrode can be evaluated by monitoring of the
OCP
which typically results in a positive shift in the OCP, inductive
of the
formation of a passive film (i.e. a surface oxide protecting the
metal from
further oxidation). The formation of porous oxide films, which can
hinder
Chapter Two Theoretical Aspects
20
but not prevent further oxidation, typically results in a decrease
in the
OCP [82, 83].
2.9.2 Linear Swap Voltage (LSV)
A LSV plot is one of the most popular DC techniques for valuation
of
corrosion rate. The induction of anodic and/or cathodic Tafel lines
for
charge transfer controlled reactions gives the corrosion current
density,
icorr, at the corrosion potential, Ecorr. This method is based on
the
electrochemical theory of corrosion processes [80].
It is performed on a metal specimen by polarizing the specimen
about
300 mV anodic ally (positive- going potential) and catholically
(negative
- going potential) from the corrosion potential, Ecorr. The
resulting current
is plotted on a logarithmic scale as shown in Fig. (2.10).The
corrosion
current, Icorr, is obtained from a Tafel plot by extrapolating the
linear
portion of the curve to Ecorr. The corrosion rate can be calculated
from
Icorr by using equation (4). The term io is a function of reaction,
the
concentration of reactants, the electrode materials, the
temperature and
surface roughness. The potential of the intersection between
anodic
reaction rate and the cathodic reaction rate is identified as
Ecorr, and the
current at this intersection is defined as icorr. The slop of the
V-I curve can
be measured. The slope is called the polarization resistance, Rp
and it is
related to the corrosion rate by:
(5)
Where are constant that are characteristic of the corroding
metal
and which can be measured or estimated.
This technique is called linear polarization
Chapter Two Theoretical Aspects
curves [78].
2.9.3 Electrochemical Impedance Spectroscopy (EIS)
Impedance electrochemical test system is one of the most popular
AC
techniques that can provide precise, error-free kinetic and
mechanistic
information by using a variety of techniques and output formats
(Nyquist
plots, Bode plots, Randles plots, etc.). This method can be used to
make
measurement in low conductivity solutions where DC techniques
are
subjected to potential-control errors. During EIS experiments, a
small
amplitude AC signal is applied to the system being studied.
Therefore, it
is a nondestructive method for the evaluation of a wide range
of
materials, including coatings, anodized films and corrosion
inhibitors.
EIS measurements can provide information regarding the kinetics
of
electrochemical corrosion system [84]. Some characteristics of
coatings,
such as corrosion behavior, porosity, solution absorption and/or
film
delimitation, can be predicted by EIS [85]. An EIS measurement,
an
alternating voltage, is applied to the corroding metal, and the
impedance,
Z is measured because both magnitude and the relative phase angles
of
Chapter Two Theoretical Aspects
22
the voltage and current have to be accounted for the impedance
is
measured for a number of frequencies that span a range from a few
mHz
to 100 kHz, and the result is complex number (the impedance) for
each of
the hundreds to thousands of frequencies that are used. In EIS, an
AC
voltage of varying frequency is applied to the sample. It is useful
to think
of the frequency as a camera shutter that can be very fast
(high
frequency) for fast reactions and very slow (low frequency) for
slow
reactions. This is the technical feature that allows EIS to get so
much
information on an electrochemical reaction in one experiment. This
is
why EIS is more useful for coatings than dc electrochemical
techniques.
EIS can quantitatively measure both resistances and capacitances in
the
electrochemical cell. A resistance corresponds to electron
transfer
reactions such as corrosion. The c